The present invention relates to magnetic resonance analysis and, more particularly, to a radiofrequency (RF) resonator for generating a substantially homogenous RF magnetic field for the purpose of magnetic resonance analysis. The present invention further relates to a method of designing the RF resonator, a magnetic resonance imaging (MRI) apparatus incorporating the RF resonator and a method of magnetic resonance analysis of an object.
MRI is a method to obtain an image representing the chemical and physical microscopic properties of materials, by utilizing a quantum mechanical phenomenon, known as Nuclear Magnetic Resonance (NMR), in which a system of spins, placed in a magnetic field resonantly absorb energy, when applied with a certain frequency.
A nucleus can experience NMR only if its nuclear spin is not zero, i.e., the nucleus has at least one unpaired nucleon. When placed in a magnetic field, a nucleus having a spin is allowed to be in a discrete set of energy levels, the number of which is determined by the spin, and the separation of which is determined by the gyromagnetic ratio of the nucleus and by the magnetic field. Under the influence of a small perturbation, manifested as an RF magnetic field, which rotates about the direction of a primary static magnetic field, the nucleus has a time dependent probability to experience a transition from one energy level to another. With a specific frequency of the rotating magnetic field, the transition probability may reach the value of unity. Hence at certain times, a transition is forced on the nucleus, although the rotating magnetic field may be of small magnitude relative to the primary static magnetic field. For an ensemble of nuclei the transitions are realized through a change in the overall magnetization.
Most MRI systems use a static magnetic field having a predetermined gradient, so that a unique magnetic field is generated at each region of the analyzed object. By detecting the NMR signal, knowing the magnetic field gradient, the position of each region of the object can be imaged. Typical MRI systems include a main magnet generating a uniform static magnetic field, whereas gradients in predetermined directions are obtained by providing additional coils which generate the desired gradients.
The rotating magnetic field in MRI systems is provided by an RF resonator, also known as RF coil, RF probe or RF antenna. The process of imaging or analyzing an object (e.g., a patient or a sample) is as follows. When pulse sequences are applied to the RF resonator, an RF radiation is emitted onto the object. According to the above principles, the RF radiation triggers NMR signals from the object from which information is obtained and subsequently used to reconstruct images and/or to analyze the object. In most MRI systems, the RF resonator is used both for transmitting the RF radiation and for detecting the resulting NMR signals from the object. RF resonators are required to generate a very uniform RF magnetic field, as any inhomogeneity in the RF magnetic field causes identical spins at different locations within the imaged object to respond differently to the RF radiation thereby distorting the image or negatively affecting the quality of the analysis. It is recognized that RF resonators which generate inhomogeneous magnetic fields also have inhomogeneous sensitivity and that RF resonators generating a weak magnetic field also detect weak NMR signals.
In addition to the homogeneity requirement, the RF magnetic field generated by the RF resonator is required to have a resonance frequency which matches the resonance frequency of the nuclear spins in the imaged sample. A known phenomenon is that once a sample is inserted into the RF magnetic field, the resonance frequency is shifted. Thus, RF resonators are typically equipped with appropriate circuitries which tune and rematch the resonance frequencies of the RF resonator and the sample. To increase signal-to-noise-ratio (SNR) and to optimize the efficiency of the system, it is also desirable that the size of the RF resonator will be comparable to the size of the sample.
Many RF resonators are presently known, and can be categorized into two groups, commonly referred to as the group of surface resonators and the group of volume resonators.
FIG. 1 illustrates a surface resonator, known as the single-loop coil [M. R. Bendall, “Surface Coil Technology”, Magnetic Resonance Imaging, ed. by C. L. Partain et al., Philadelphia, 1988]. The single-loop coil is a planar current-loop, which is sensitive to RF fields, designated herein by B1, in the direction perpendicular to the loop surface. The single-loop coil is characterized by a high SNR, however, its effective homogenous RF field is only near the surface, or the plane, of the current-loop. Thus, the single-loop coil is highly inhomogeneous.
FIG. 2 illustrates another surface RF resonator, known as the phased array coil [Sodickson D K and Manning W J., “Simultaneous acquisition of spatial harmonics (SMASH): ultra-fast imaging with RF coil arrays”, Proc. Fifth Scientific Meeting of the International Society for Magnetic Resonance in Medicine, 1817 (1997); Klass P. Pressmann, et al., “SENSE: Sensitivity Encoding for fast MRI”, Magnetic Resonance in Medicine, 42:952, 1999]. The phased array coil is an array of a number of single-loop coils, where the phases of the single-loop coils are designed so that the matrix representing their signals is can be diagonalized. The volume of interest of the phased array coil is the combined near-surface of each single unit of the array.
As opposed to the surface resonators, where the effective imaged region is a surface (i.e., two-dimensional space), in the group of volume resonators the effective imaged region is a volume (three-dimensional space). FIG. 3 illustrates a volume RF resonator known as a saddle coil [D. W. Alderman et al., J. Magn. Reson. 36, 447, 1979]. The saddle coil is made of a combination of two current-loops, wrapped around a lateral surface area of a cylinder, so that the magnetic fields of the two current-loops combine. In some saddle coils each loop is made a multi-loop structure in a spiral manner.
FIG. 4 illustrates another volume RF resonator known as the multi-turn solenoid [D. I. Hoult and P. C. Lauterbour, J., “The Sensitivity of Zeugmatographic Experiment Involving Human Samples” Magn. Reson, 34:425, 1979 ]. The multi-turn solenoid is a structure which generates a very homogenous magnetic field within a cylindrical volume. The multi-turn solenoid is rarely used in MRI because the magnetic field is parallel to the cylinder axis and is therefore orthogonal to the symmetry of clinical systems where the RF field is to be orthogonal to the longitudinally oriented main static magnetic field.
FIG. 5 illustrate an additional volume RF resonator known as the single-turn solenoid [J. P. Hornak, et al., “Elementary Single Turn Solenoids Used in the Transmitter and Receiver in Magnetic Resonance Imaging”, Magn. Reson. Imag., 5:233-237, (1987)]. The single-turn solenoid formed by a broad sheet instead of a multi-turn structure. Conceptually, it is very similar to the conventional solenoid, thus, it suffers from similar limitations as the multi-turn solenoid and it is therefore used only for imaging of particular samples, such as the breast and forearm. One type of single-turn solenoid is the perforated single-turn solenoid which is a single-turn solenoid with one or more perforations in its cylinder. The perforated single-turn solenoid is suitable for imaging of head, knee, elbow, wrist and shoulder.
FIG. 6 illustrates the most common volume RF resonator for clinical imaging, known as the birdcage resonator or the birdcage coil [C. E. Hayse et al., “An Efficient Highly Homogeneous Radiofrequency Coil for Whole-Body NMR Imaging at 1.5 T”, J. Magn. Reson., 63:622-628, 1985; J. Tropp, “The theory of the bird-cage resonator”, J. Magn. Reson. 1989; 82, 51-62; T. A. Riauka el al., “A numerical approach to non-circular birdcage RF coil optimization: Verification with a fourth-order coil”, Magnetic Resonance in Medicine, 41:1180-1188, 1999; Jianming J. “Electromagnetic analysis and design in magnetic resonance imaging”, CRC Press, New York, 1999]. The birdcage coil is a very homogenous resonator which is formed from a number of equally spaced conductors on a cylindrical surface. It is common to refer to conductors which are longitudinally oriented with respect to the symmetry axis of the cylinder as “legs” or “rungs”, and to conductors which are transversally oriented as “end-rings”. Capacitors are located on the legs, on the end-rings or both on the legs and the end-rings.
Currents flowing in the legs and end-rings obey an eigenvalue equation which is characterized by a set of eigenvectors, also known as current-modes. Each current-mode corresponds to a set of currents flowing in the legs, and is associated to an eigenvalue which is related to one possible solution of resonant frequency. Specifically, a birdcage coil with N legs has N resonant frequencies and N current-modes. For a linear birdcage, one special current-mode, has a sinusoidal current distribution among the legs of each side along the circumference. In this mode, the magnetic field inside the resonator is very homogenous. In is known that the homogeneity level of the field is proportional to the number of legs, where an infinite number of legs incorporate a desired mode corresponding to an RF field having a perfect homogeneity.
However, as the number of current-modes increases, so does the complexity of the birdcage coil design. The undesired current-modes of the birdcage coil, except for intrinsically orthogonal cosine mode, reduce the field homogeneity and/or the transmission and detection power. Thus, when designing a birdcage coil, one needs to eliminate all current-modes other then the desired mode to ensure a homogenous field. This is typically done by designing a birdcage coil where the desired mode resonant frequency is significantly spaced apart from all the other frequencies. An additional factor which is to be considered when selecting the number of birdcage coil legs is the physiological effect of a large number of legs on the patient which may become claustrophobic.
An inherent limitation of the birdcage coil is that in the design process, both the magnetic field characteristic and the resonance characteristic of the birdcage coil are to be simultaneously calculated, as these two characteristics are entangled. Any change in the number of legs and separation therebetween and the location of the capacitors and the type thereof alters both the RF field lines and the resonant frequency of the coil.
Moreover, the simulations preceding the manufacturing of the birdcage coil exhibits a very homogenous RF magnetic field. In practice, however, once a sample is inserted into the coil, the resonance frequency and the tuning impedance of the coil are shifted. Although this effect can be approximated during simulation, the validity of such approximation is very limited due to the various sizes, structures and orientations that a biological organ may exhibit when placed in the birdcage coil, contrary to the spherical or cylindrical uniform sample that is typically used for simulations. Thus, the resonance frequency and the tuning impedance must be retuned by an arrangement of tunable capacitors, which has to be electrically connected to the birdcage coil. Theoretically, N tunable capacitors can correct some of the sample effects, but for practical reasons only a few capacitors are used. The additional capacitors break the symmetry of the birdcage coil, result in loosing field homogeneity and introduce non-zero contributions from one or more undesired current-modes. In realistic birdcage coils, the inhomogeneity of the RF field may approach 15-20%.
Additional prior art of relevance is a volume RF resonator known as the Litz coil, disclosed in U.S. Pat. No. 6,060,882 and illustrated in FIG. 7. The coil is based on Litz foil conductor, which includes multiple and parallel Litz wires with interwoven sub-routes from a first node to a second node and insulated crossovers forming well-defined flux sub-windows. The structure of the coil results in multiple current routes each contribute to the RF magnetic field, B1, leaving a central flux window centered on the B1 axis. An identical semi-coil, formed on the opposite side of the sample around completes the coil. The two semi-coils are electrically connected in parallel. This coil solves some of the problems associated with the birdcage coil, however its design and manufacturing is still rather complicated. Specifically, the Litz coil efficiency depends on the number of braids and on their thickness. Hence, for a sufficiently efficient Litz coil the braiding is very complex and ultra thin.
Also of prior art of relevance is the so-called “Slotted Tube Resonator” [H. J. Schneider and P. Dullenkopf, “Slotted Tube Resonator: A new NMR probe head at high observing frequencies”, Rev. Sci. Instrum., 48:68-73, 1977]. This resonator is composed of a conducting tube which is cut lengthwise, thereby forming a strip line which consists of two arched conductors. The slotted tube resonator is coupled to an electronic circuitry, which includes two independently operating capacitors, a series capacitor for matching the resonant characteristics of the resonator and a shunting capacitor for tuning the impedance of the resonator.
An improved type of the slotted tube resonator [D. H. Hong et al., “Whole Body Slotted Tube Resonator for Proton NMR Imaging at 2.0 Tesla”, Magn. Res. Imag., 5:239-243, 1987;], incorporates a coupling sheet, positioned externally to the tube arches, for linking the resonator with the transmitter and/or the receiver of the imaging system. In addition, the improved slotted tube resonator includes additional capacitors forming a capacity coupling between the two tube arches.
One application of the slotted tube resonator, designed specifically for analyzing hearts of rabbits of different sizes, is described in G. J. Kost, “A Cylindrical-Window NMR Probe with Extended Tuning Range Studies of the Developing Heart”, J. Magn. Res. 82:238-252, 1989. This coil is cylindrical and it is formed with a window in the cylinder wall. Unlike the multi- and single-turn solenoids, the produced magnetic field of the cylindrical windowed coil is directed perpendicularly to the symmetry axis of the cylinder. This coil, however, is suitable only for non-imaging applications of small sized samples.
Other slotted tube resonators are described in S. Bobroff and M. J. McCarthy, “Variations on the Slotted-Tube Resonator: Rectangular and Elliptical Coils”, Magn. Res. Imag., 17:783-789, 1999; M. K. Murphy et al., “A Comparison of Three Radiofrequency Coils for NMR Studies of Conductive Samples”, Magn. Res. in Med., 12:382-389, 1989; T. Sphicopoulos and F. Gardiol, “Slotted Tube Cavity: a Compact Resonator With Empty Core”, IEE Proceedings, 134:405-410; A. Darrasse et al., “The slotted cylinder: an efficient probe for NMR imaging, Magnetic Resonance in Medicine, 2(1):20-8, 1985; H. J. Schneider and P. Dullenkopf, “Crossed Slotted Tube Resonator: A new Double resonance NMR probehead”, Rev. Sci. Instrum., 48:832-834, 1977; D. W. Alderman and D. M. Grant, “An Efficient Decoupler Coil Design which Reduces Heating in Conductive Samples in Superconducting Spectrometers”, J. Magn. Res., 36:447-451, 1979; S. Crozier et al., “In Vivo Localized 1H NMR Spectroscopy at 11.7 Tesla”, J. Magn. Res., 94:123-132, 1991; C. Ranasinghage et al., “Resonator Coils for magnetic resonance imaging at 6 MHz”, Med. Phys. 15:235-240, 1988; and T. A. Gross et al., “Radiofrequency Resonators for High-Field Imaging and Double-Resonance Spectroscopy”, J. Magn. Res., 62:87-98, 1985.
It will be appreciated that, all of the prior art slotted tube resonators suffer from two crucial limitations: (i) the slotted tube resonator fails to provide a mechanism for balancing the RF field homogeneity once it has deviated from its original design due to sample-field interactions; and (ii) the geometrical configuration of the slotted tube resonator is such that the generated RF magnetic field has a linear polarization. An intrinsic limitation of the linear polarization is the power losses, which are explained by wasted components in the mathematical expansion of the linear polarization. In a theoretical study of current distributions and field uniformity in saddle coils [J. W. Carlson, “Currents and Fields of Thin Conductors in RF Saddle Coils”, Magn. Res. in Med., 3:778-790, (1986)] an optimal geometry for circular polarized coil has been calculated.
In this respect, one coil for generating a circular polarized magnetic field is of the birdcage type described above, also known as the quadrature birdcage coil, in which intrinsically orthogonal sine and cosine modes are incorporated. A particularly interesting quadrature birdcage coil is described in an article by H. Barfuss et al., entitled “In Vivo Magnetic Resonance Imaging and Spectroscopy of Humans with a 4 T Whole-body Magnet”, published in NMR in Biomedicine, 3:31-45. This birdcage coil consists of four longitudinal and two transverse copper foils, interconnected via a system of capacitors. The coil is further equipped with additional rotary differential capacitors for providing continuous distribution between the tuning and matching capacitors, thereby allowing variation of the quality factor of the coil. However, this birdcage coil only partially addresses to problems associated with effects of interaction between the magnetic field and the imaged object. Specifically, this birdcage coil, although successfully providing a sophisticated circuitry for tuning the resonance characteristics and matching the impedance, fails to provide a satisfactory mechanism for correcting inhomogeneity in the magnetic field once a sample is placed therein.
There is thus a widely recognized need for, and it would be highly advantageous to have, an RF resonator and a method of designing the same, devoid of the above limitations.